At present, approximately 300 000 patients worldwide benefit each year from a valve prosthesis replacing one or more of their heart valves damaged either by infectious disease or by a degenerative process linked to aging.
There are two major families of prosthetic heart valves:                valve prostheses of biological origin, known as bioprostheses, which are either removed from the animal and treated chemically or constructed from biological tissues on the model of a natural valve;        mechanical valve prostheses, which are devices unrelated to the shape of a natural valve and manufactured from biologically compatible man-made materials resistant to wear.        
Because of their anatomical configuration and their physiological mode of operation, bioprostheses offer biological performance comparable with that of a natural heart valve because they respect the natural structure of the flow of blood through the cavities of the heart and the aorta.
This feature of bioprostheses saves patients lifelong anticoagulant treatment, which eliminates the risk of hemorrhage following long-term use of these drugs and therefore improves the quality of life of these patients.
Thus patients may even forget that they have an artificial heart valve.
Moreover, it is necessary to note that bioprostheses do not cause any acoustic disturbance, which also helps patients to forget that they have an artificial heart valve.
These bioprostheses have a limited service life, however, because of unavoidable calcification over time, which imposes their replacement after a period of about ten years on average. Once started, this calcification accelerates and destroys the valve, with the consequence of progressive deterioration of the valve function and its repercussions on the heart muscle. This calcification occurs more quickly in young persons than in elderly persons, which limits the field of application of bioprostheses to persons aged 65 or more or persons whose life expectancy is less than the service life of the bioprosthesis.
In France the life expectancy at age 65 is 17.7 years for men and 21.7 years for women and replacing a defective heart valve is regarded as major surgery that is accompanied, beyond the age of 75, by a high mortality rate. To this risk is added the discomfort, at this age, of major surgery.
In contrast to bioprostheses, mechanical artificial valve devices are not degradable and have a service life exceeding the human lifetime. On the other hand, because their geometry departs very considerably from the natural model and because of their non-physiological mode of operation, these mechanical valves generate on each heart beat disturbances to the flow of blood in the form of turbulence, areas of recirculation, vortices, shearing of the blood cells and slowing or stasis of the flow over the parts of the mechanical device that are poorly swept by the blood flow, notably the articular areas.
These disturbances to the flow increase the time of contact of the blood cells with and the intensity of the reactions of active proteins on the prosthetic materials constituting these devices. Any foreign material in contact with the blood inherently stimulates the coagulation process. There results from interaction between the disturbances to the flow and the non-biological materials:                adhesion to the surface of these materials of active proteins and blood platelets,        activation of coagulation, and        formation on the surfaces of organized bloodclots.        
This powerful biological phenomenon is the same as that which governs the physiological process of healing the internal wall of the blood vessels. It prevents blood leaking out of the circulatory system. It is therefore indispensable to life and difficult to counteract.
However, not only can coagulation deposits impede the mechanical function of the valve on blood circulation, which puts the life of the patient at risk, but these deposits can also migrate in the circulatory system (embolisms), most often in the cerebral circulatory system, and lead to neurological problems, often with invaliding repercussions.
To these coagulation phenomena is added trauma of the red cells, repeated on each cardiac cycle, which shortens their life (hemolysis) and leads to chronic inflammatory reaction of the entire organism. This reaction itself tends to increase the coagulability of the blood, which increases the probability of coagulation incidents.
Thus thrombosis generates thrombosis and creates a self-sustaining chronic illness.
To remedy this drawback, any patient with a mechanical artificial valve device must be protected for their entire life by anticoagulation treatment, with the inherent risk either of hemorrhage in the event of an overdose or of thrombosis/embolism incidents in the event of underdosing.
Since the beginning of the sixties, a number of generations of mechanical heart valves have been designed to reduce the disturbance to the blood flow that these devices generate, so as to reduce the risk of coagulation: firstly, valve prostheses consisting of a ball floating in a cage (STARR-EDWARD), then, at the beginning of the seventies, second generation prostheses consisting of a pivoting disk (BJÖRK-SHILEY) and then, ten years later, third generation prostheses with two flaps and lateral openings (ST-JUDE MEDICAL). This third generation is that most widely used at present and produced in different forms by a number of manufacturers.
Despite these improvements, third generation valves still cause blood trauma and still cannot function in man without anticoagulant drugs. On the other hand, thanks to more than 40 years of clinical experience, the anticoagulation treatment is now well codified.
Patients with a mechanical valve in the aortic position must maintain their blood coagulability (as measured by a standardized biological method known as the international normalized ratio (INR) method) at a level at least two and a half times greater than the physiological value (INR 2.5).
Patients with a mechanical valve in the mitral position must maintain their blood coagulability at a level at least three and a half times the physiological value (INR 3.5).
This difference of in the “harmfulness” of mechanical prostheses between the aortic position and the mitral position is a result of the fact that blood flows more slowly through the mitral orifice than through the aortic orifice. The time to fill the heart through the mitral valve (typically of the order of 450 milliseconds at 70 beats per minute) is longer than the time to eject the blood through the aorta (typically of the order of 300 milliseconds). The blood is therefore in contact with a prosthetic valve in the mitral position for longer, which enables coagulation processes to complete.
Moreover, mitral valves are larger, and so the areas of foreign materials exposed to biological deposits are greater. Thus it has been established that the risk of thrombosis/embolism complications in patients with mechanical heart valves is twice as high for the mitral position as for the aortic position.
For large numbers of patients with mechanical heart valves, the average rate of coagulation incidents under current medical practice is statistically less than 3% per annum and per patient and the rate of hemorrhage is less than 4% per annum per patient.
This state of the art data provides a benchmark for clinicians to evaluate the thrombogenetic potential of a new mechanical heart valve during probationary testing in man and is decisive for obtaining authorization to place it on the market. A rate of thrombosis/embolism or hemorrhage complications greater than 3-4% will lead to rejection of the product by the medical community and refusal of licenses.
As long as the anticoagulant protection is correctly administered, millions of patients with mechanical heart valves worldwide can nevertheless now live under acceptable conditions. These patients, who were previously condemned to die within a short time, now live on for many years. However, their life expectancy, given the risk of thrombosis/embolism and hemorrhage, remains significantly less than that of persons of the same age without a heart valve.
The imperative requirement for anticoagulation protection for all patients with mechanical heart valves is particularly dramatic in countries where the medical infrastructure does not provide satisfactory monitoring of anticoagulation treatment. In those countries, valve disease is becoming endemic and is more likely to affect younger persons, women and the mitral valves. For example, in India several million children under the age of 15 need a prosthetic valve replacement. These young persons are poor candidates for biological type valves because of the calcification problems referred to above. Mechanical heart valves are therefore more willingly employed but are accompanied by a rate of dysfunction through coagulation that is very much higher than is observed in developed countries, and this major risk is restricting their use. In these countries the thrombogenetic nature of mechanical heart valves represents a public health problem and illustrates the need for better performing products whose use would impose fewer constraints.
It should be noted that even if the anticoagulation treatment is followed correctly, the rate of complication remains of concern even in countries in which the medical infrastructures are adequate. Statistically speaking, over a period of 10 years, one in two mechanical heart valve patients will have experienced a serious complication necessitating hospitalization, either because of a coagulation incident or because of a hemorrhage.
Mechanical heart valve designers are therefore seeking to improve the hydrodynamic performance and the mode of operation of these devices to reduce the disturbances that they cause in the flow of blood and thereby to eliminate or at least to reduce the doses of anticoagulation drugs necessary to prevent these complications.
There is known from U.S. Pat. No. 6,395,024 a mechanical prosthetic heart valve that includes a ring with an interior peripheral surface centered on an axis and three flaps disposed in the vicinity of the interior peripheral surface of the ring. These three flaps are adapted to pivot between a closed position preventing blood from flowing through the valve and an open position in which blood flows axially through the valve.
The ring has an edge called the downstream edge that connects the interior peripheral surface to an exterior peripheral surface and is on the downstream side of the flow and three crenelations or protuberances that extend axially in the downstream direction from this edge.
Each flap has a central part provided with two lateral wings each of which cooperates with respective means for guiding rotation of the valve provided on the interior surfaces of two consecutive crenelations. The space in which each lateral wing of a flap pivots is called the pivoting space.
Also, two windows are formed symmetrically in each crenelation.
Each window enables satisfactory rinsing of the external face of the lateral wings of the flaps by the retrograde flow.
Accordingly, when the valve is installed in the mitral position, this external face can be swept by the flow of blood circulating from the ventricle toward the aorta. Thus, thanks to this feature, all risk of biological deposits at this location is eliminated.
Likewise, when the valve is installed in the aortic position, the reflux of blood through these windows into the aortic sinuses when the valve is closed can rinse the external face of the lateral wings, preventing a volume of blood from being trapped in the pivoting spaces of the flap.
To perfect this protection against stagnation of blood in the pivoting spaces, an additional feature has been provided: the lower edge of the windows described above forms with the leading edge of the lateral wings of the flaps, when the latter are open, a second opening having a triangular sluice shape. This second opening (called a “cleft”) is “dynamic” in the sense that the area of the orifice formed in this way increases progressively as the flap moves from the closed position to the open position. It allows direct passage to the exterior of the flaps of blood conveyed by the anterograde flow and assures additional sweeping of the leading edge and the external face of the wings of the flaps.
The Applicant has nevertheless noticed from implantation in animals that the effect of this additional feature on the flow of blood is not the same in the mitral position as in the aortic position.
The above feature proved to be efficacious on a large number of animals implanted with the valve in the mitral position and left for many months without anticoagulation protection, whereas this was not the case with animals in which the same valve had been implanted in the aortic position.
In the mitral position, during ventricular filling, low-pressure blood can flow through the second openings (“clefts”) from the interior of the valve toward the exterior in the pivoting spaces of the flaps and rinse these critical pivoting spaces.
However, during ventricular ejection, the blood pressure generated by the heart on a valve implanted in the aortic position is ten times greater than the blood pressure that is exerted via a valve implanted in the mitral position.
Because aortic valves are smaller than mitral valves and the clefts are therefore much narrower, in the aortic position the effect of rinsing is to create, on each heart beat, powerful lateral “jets” that go beyond the intended objective of rinsing to the point of causing trauma to the blood cells.
The trauma threshold recognized in the prior art in this connection is situated at a force of around 150 dynes/cm2 for blood platelets and 1000 dynes/cm2 for red cells. Beyond these values, blood components are sheared and the blood platelets release their coagulating agents, which can cause coagulation complications.
Thus clefts that are efficacious in the mitral position to prevent slowing of the blood in the pivoting spaces are of no utility and potentially dangerous in the aortic position.
Clinical experience has shown that the articulation areas of a mechanical heart valve are the areas most exposed to coagulation phenomena.
Unfortunately, as a heart valve assures on each heart beat a function that is vital to the circulation of the blood, the specifications imposed by functional safety imperatives have to take priority over coagulation problems.
Thus the flap articulation mechanism imposes a geometry that is not favorable to a good blood flow structure in the pivoting spaces. It generates shear and microturbulence in the immediate vicinity of surfaces that are relatively poorly swept by the blood stream.
The amplitude of this phenomenon is linked to the number of articulation areas of the valve. It is therefore greater for a heart valve with three flaps that comprises six pivoting spaces than for a heart valve with two flaps, which has only four such spaces.
Because of this, the advantages of the mechanical heart valve with three flaps where resistance to coagulation complications is concerned are found to be greatly reduced if specific devices are not fitted.
Patients who need a prosthetic heart valve wish to be operated on only once and to be protected from coagulation complications that can arise if foreign bodies are present in the circulatory system.
Unfortunately, to prevent coagulation deposits forming, patients are obliged to take anticoagulant drugs throughout their life, which is a constraint, and long-term use of such drugs is liable to induce hemorrhage complications.